The present invention relates to the art of magnetic resonance imaging of moving substances and velocity determination by phase mapping. It finds particular application in conjunction with measuring blood flow and will be described with particular reference thereto. However, it is to be appreciated that the invention is also applicable to imaging or phase mapping other anatomical fluid flows, non-biological fluids, multi-component systems in which one gaseous, fluid, semi-fluid, solid, or other component moves relative to other components, or the like.
Heretofore, it has been recognized that the spin-echo magnetic resonance imaging signal from material flowing through a field gradient experiences a phase shift. Because the phase shift varies linearly with velocity, the spin-echo imaging and phase mapping have been used to measure velocities of slow moving, uniform flows, including blood flows in vivo. However, varying flow rates cause corresponding variations in the phase shift which interferes with the phase encoding. These phase variations result in noise which extends from the blood vessel across the image in the phase encoding direction. This is particularly serious at high flow velocities at which there are frequently small velocity variations and eddy currents which cause significant phase shifts. This renders the prior art techniques unsuited to phase mapping high blood flow velocities such as major blood flows adjacent the heart. High blood flow rates of clinical interest occur in the coronary, carotid, and pulmonary arteries and the aorta.
Several techniques have been devised for phase mapping slower, relatively constant blood flows. In one such technique, the gradients are velocity encoded. The pixel phases are calculated on one image obtained using the velocity encoded gradients. See P. R. Moran, "A Flow Velocity Zeugmatographic Interlace for NMR Imaging in Humans," Magnetic Resonance Imaging, volume 1, pages 197-203, 1982. In another technique described in D. J. Bryant, J. A. Payne, D. N. Firmin, and D. B. Longmore, "Measurement of Flow with NMR Imaging Using a Gradient Pulse and Phase Difference Technique", in the Journal of Computer Assisted Tomography, volume 8, number 4, pages 588-593, 1984, two images are obtained. One is obtained with a standard sequence and the other is obtained with velocity encoded gradients. The phase difference between the two is then mapped. In a technique of T. W. Redpath, D. G. Norris, R. A. Jones, and J. S. Hutchison, "A New Method of NMR Flow Imaging", Phys. Med. Biol. volume 24, number 7, pages 891-898, 1984, the velocity information is extracted pixel by pixel. Several images, each with one of a plurality of encoded gradients, are collected. A Fourier transform is performed on the pixel phases to extract the velocity information.
Two-echo Carr-Purcell sequences have also been used to map blood flows. However, the Carr-Purcell sequences make no determination of the flow velocity. The second echo of the Carr-Purcell sequence refocuses the flow related phase shift information. That is, the flow related phase shift acquired by the material flowing in the read gradient direction on the first echo is refocused on the second echo. This refocusing provides increased intraluminal signal intensity with even-echo rephasing. See K. J. Packer, "The Study of Slow Coherent Molecular Motion by Pulsed Nuclear Magnetic Resonance", Molecular Physics, volume 17, number 4, pages 355-368, 1969 and V. Waluch and W. G. Bradley, "NMR Even Echo Rephasing in Slow Laminar Flow", Journal of Computer Assisted Tomography, volume 8, number 4, pages 594-598, 1984.
Conventional slice selection and frequency encoding both use spin or field gradient echoes. The phase shift at each pixel is dependent on the flow in both the read and the slice select direction. Although the relative phase shift at each pixel may be rendered sensitive to flow in only a single direction, the high sensitivity to flow velocity variations remains. Commonly, a phase shift of greater than 20 degrees for each centimeter per second of velocity change is experienced.
The velocity sensitivity has two adverse effects on the resultant image. First, magnetic resonance signals are phase dependent on the varying phase-encode field gradient. The two dimensional Fourier transform method of spatially mapping the magnetic resonance signals identifies signal components arising from different positions along the phase-encode gradient. As the gradient is incremented to collect a plurality of views, the signal arising from a given position changes in phase by an amount proportional to its position. Spatial resolution is improved by the Fourier transformation of the signals acquired over a plurality of views or increments of the phase encoded gradient, e.g. 256 views.
The phase dependence of the magnetic resonance signal on velocity introduces an extra term in the phase encoding process. When the velocity dependent phase term is constant from view to view, the extra term has no effect on the spatial resolution. However, when the velocity changes from view to view, such as from turbulence or minor irregularities in heart rate or ejection volume, the associated phase changes are interpreted as noise when decoding the spatial position of the resonance signal component. The available signal intensity is dispersed through the image in the phase encoding dimension or direction. This prevents signal phase, hence, velocity determination.
The second effect of the velocity sensitivity is a reduction in the signal intensity from vessels flowing in the read direction by phase cancellation through a flow profile. For example, laminar flow through a circular pipe lying within the selected slice has a flow profile which is parabolic. The maximum flow velocity is at the center of the pipe, while the flow velocity is zero at the boundary with the wall. When the pipe and the fluid flowing therein is imaged through a longitudinal slice, one resultant resonance signal represents the projection or integral of the signal components from columns of incremental volume elements extending across the pipe. The component from each incremental volume element has the same intensity but each has a phase which is proportional to the local velocity. Summing or projecting the phases in the column across the pipe sums components from incremental volume elements which have the generally parabolic range of velocities. This summation of signal components across the pipe leads to signal attenuation by phase cancellation. As the phase sensitivity to velocity increases, the phase cancellation becomes greater. In a typical spin-echo sequence which is imaging velocities in the read direction above 15 centimeters per second, attenuations of over 75% have been encountered. In actual measurement of blood flow, the attentuation is lower attributable to the non-Newtonian nature of blood. Blood travels through the body with a flow which approximates plug flow more closely than laminar flow.
Another disadvantage of spin-echo sequences for blood measurement arises from the use of a single 180 degree pulse within each cycle of the sequence. The 180 pulse generates the echo and refocuses the applied gradients as well as the static magnetic field inhomogeneity. If the refocusing pulse is slice selected, excited material flowing oblique to or perpendicular to the selected slice may move out of the selected slice in the time interval between application of the excitation and refocusing pulses. Hence, the excited material does not experience the refocusing pulse and severe signal loss is observed. However, this disadvantage may be overcome with broadband refocusing pulses which refocus excited material however it has moved since excitation.
However, broadband refocusing pulses introduce another problem in blood and other materials with relatively long relaxation times. Material outside the selected slice is subject to a series of the broadband refocusing pulses arising from repetition of the imaging sequence. This has the effect of inverting Z-magnetization at every repetition of the imaging sequence, typically at least every second. The inversion pulses then cause a progressive saturation of the material outside the slice until a semi-saturated steady state polarization is reached. When this material moves into the slice, it produces a signal proportional to this polarization. Partial saturation then produces a corresponding reduction in signal intensity. For example, when imaging blood with a 20 ms spin-echo sequence repeated every one second, the intensity is reduced approximately 45% relative to fully polarized blood.
The present invention contemplates a technique which overcomes the above referenced problems and others to map even high velocity and non-constant blood flows.